An estimated 4.8 million Americans suffer from congestive heart failure (CHF). Many of these persons are unresponsive to pharmacological intervention and could benefit from a heart transplant. As a result of the current shortage of donor hearts, implantable blood pumps have gradually evolved into a viable treatment option for these persons. Over the past 30 years, many devices have been developed that either replace the entire heart or assist the heart.
In a diseased state, one or both of the ventricles of the heart can become greatly weakened to an extent that mechanical intervention is needed to keep a patient alive. In some instances, the entire heart is removed and replaced with a total artificial heart while in other cases a device that assists the heart is used. A blood pump used for assist is commonly referred to as a ventricular assist device or VAD.
Although either of the ventricles of the human heart may function in a weakened state, it is the left ventricle that is primarily treated for insufficient pumping. Normally, blood enters the left ventricle through the mitral valve and, during heart systole, the blood is ejected through the aortic valve and into the aorta by the squeezing action of the left ventricle. To assist a failing left ventricle, a VAD is typically attached between the apex of the left ventricle and the thoracic aorta. In this way, blood entering the left ventricle may either be ejected through the aortic valve by the ventricle or pass through the VAD.
Ventricular assistance has been performed by a variety of blood pump designs. The majority of the early VADs pumped blood in a pulsatile manner. In this case, the VAD has an internal sac situated between two heart valves. The sac is typically allowed to passively fill with blood, then the VAD pumping mechanism squeezes the sac, ejecting the blood into the patient's aorta. Patents such as U.S. Pat. Nos. 5,599,173, 5,980,448 and 4,023,468 teach devices that move blood in this manner. These pulsatile VADs are typically large and can only be used as an implantable treatment option for patients with a large body surface area, such as a large man. In addition, reliability issues exist due to the heart valves that are required for pulsatile pumping.
More recently, continuous flow pumps are being developed to address the size and reliability requirements. VADs such as described in U.S. Pat. Nos. 5,947,892, 5,613,935, 6,074,180, 5,928,131, and 6,050,975 operate in this fashion. These pumps are smaller than their pulsatile counterparts and can be inherently more reliable. VADs having a magnetically suspended rotor, such as those described in U.S. Pat. Nos. 5,928,131, 6,050,975, 6,074,180, and 6,302,661, can have only one moving part, the pump rotor.
The connection of a VAD to the human anatomy typically requires two conduits or tubes: an inflow conduit to allow blood into the VAD and an outflow conduit to pass the pressurized blood exiting the VAD to the patient's aorta. Although these necessary components are required for nearly all VAD implantations, they are typically overlooked as components that affect VAD performance, in particular their propensity for inducing blood damage.
Previous work has disclosed various connection schemes, such as described in U.S. Pat. Nos. 4,086,665 and 5,511,958, while others have dealt with the conduits as part of an assist device design such as described in U.S. Pat. Nos. 6,346,071 and 6,319,231. Although the amount of published literature is limited, studies examining the affect of inflow conduit tip shape have shown its importance, for example “Novel ventricular apical cannula in vitro evaluation using transparent, compliant ventricular casts”, by Curtis et al. This work has clearly shown that it is possible to substantially improve the flow pattern immediately adjacent to the conduit tip through improved tip design.
For a continuous flow pump, the inflow conduit is subjected to additional pressure swings not present for pulsatile pumps. The passive filling of pulsatile pumps maintains an inflow conduit pressure that closely matches the pressure in the left ventricle. This is due to the valve that is typically positioned between the inflow conduit and the blood pump chamber For continuous flow pumps, no valve is present between the inflow conduit and blood pump. Consequently, any pressures generated upstream of the blood pump impeller blades are reflected within the inflow conduit and the left ventricle.
A low-flow and low-pressure condition is possible for continuous flow pumps if the pump continues to operate at a given speed and the left ventricle has an insufficient volume of blood available for pumping. In this instance, the pressure in the left ventricle and the inflow conduit drops quickly and both are subjected to a negative pressure. This negative pressure can cause the ventricle and/or inflow conduit to collapse. If the pump speed is not reduced, the collapse can continue and the negative pressure can cause gases to be drawn out of the blood. This phenomenon is known as cavitation and has been shown to cause blood damage.
Most of the currently available continuous flow pumps use manual control to govern the assist level of the pump. To avoid collapse and low pressures mentioned above, the pump assist level is typically set to a low level in order to avoid damaging the blood. However, this method of pump control can result in an insufficient level of assist for the patient. A preferred method for controlling a blood pump can be to measure some physiologic parameter, such as left ventricular pressure or heart size, and use that parameter to govern pump operation. Such a control scheme has been described in U.S. Pat. No. 5,928,131 for continuous flow pumps.
A pressure sensing scheme has been disclosed in U.S. Pat. No. 6,171,253 in which geometric changes in a blood carrying conduit are sensed and used as an indication of blood pressure within the blood carrying conduit. A section of the blood carrying conduit is somewhat flattened to provide a focal point for any pressure-induced shape change of the conduit. A set of strain gauges can then be used to measure the flexure of the flattened portion when pressure changes in the conduit occur. In another embodiment, the flexure can also be measured using an optical sensor that detects how far the flattened portion has moved. Yet another embodiment uses two fluid compartments to measure changes in the geometry in the flattened region of the conduit. One of the fluid compartments is partially bounded by the flattened portion of the conduit. The second compartment is situated adjacent to the first compartment with a flexible diaphragm between the two. Monitoring the pressure in the second compartment allows indirect sensing of the pressure with in the blood carrying conduit, since changes in the blood pressure within the conduit produce pressure changes within the first fluid compartment and consequently the second fluid compartment.
Several inventions have been disclosed which obviate the need for an inflow conduit, U.S. Pat. Nos. 4,944,078 and 5,507,629. This VAD is intended for placement within the left ventricular volume, thus occupying the space that is normally filled with blood. This placement negates the need for an inflow conduit, thus removing an implanted component that may cause blood damage or have other possible failure modes. However, the pump inlet is very close to the heart tissue and it is likely that negative pressure spikes could quickly develop for this VAD. The inflow conduits normally used have a portion of vascular graft incorporated within their design. Strictly speaking, collapse of an inflow conduit is undesirable, but with the VAD in U.S. Pat. No. 5,507,629 no compliant graft is positioned between the heart tissue and the pump. Consequently, any negative pressure spike will be solely imposed upon the heart tissue, making the likelihood of drawing heart tissue into the pump much greater.
U.S. Pat. No. 5,599,173 describes the structuring of the inflow channel of a VAD to produce beneficial flow patterns in the blood sac of a VAD. The inflow channel of the above invention is tapered to direct the blood entering the VAD toward the annular wall of the sac. This produces uniform flow that minimizes the chance of thrombus formation on the blood-contacting surface of the sac.
A custom inflow conduit is needed that conforms to the patient's anatomy, allows the blood entering the blood pump to do so in a stagnation- and recirculation-free manner, and prevents suction-induced collapse.